Biocompatible, implantable hearing aid microactuator

ABSTRACT

A biocompatible, implantable microactuator (82) for a fully implantable hearing aid system includes a hollow body (84) that has an open first end (88) and, open first and second faces (94a and 94b). Flexible diaphragms (92, 96a and 96b), respectively covering the end (88) and faces (94a and 94b), hermetically seal the body (84). An incompressible liquid (98) fills the body (84). Transducers (102), provided by laminated, stress-biased unimorphs (32 or 62) that are mechanically coupled to the flexible diaphragms (96a and 96b), deflect the diaphragms (96a and 96b) in response to an electrical driving signal. Deflections of the diaphragms (96a and 96b) are coupled by the liquid (98) to the first flexible diaphragm (92). The unimorphs (32 or 62) include a layer of biocompatible metal (36 or 66-68) deposited on one side of a biocompatible piezoelectric ceramic plate (34 or 64) to stress-bias the plate (34 or 64). A thin, biocompatible electrode (44 or 72) coats the other side of the plate (34 or 64).

This is a division of application Ser. No. 08/896,969 now U.S. Pat. No.5,977,689 filed Jul. 18, 1997, that also claims the benefit of U.S.Provisional Patent Application Ser. No. 60/022,182 filed on Jul. 19,1996.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to the field of implantable biocompatibletransducers, particularly those useful for a microactuator included in afully implantable hearing aid system.

2. Description of the Prior Art

Patent Cooperation Treaty ("PCT") patent application Ser. No.PCT/US96/15087 filed Sep. 19, 1996, entitled "Implantable Hearing Aid"("the 15087 PCT Patent Application") describes an implantable hearingaid system which uses a tiny implantable microactuator. A PCT patentapplication Ser. No. PCT/US97/02323 entitled "Improved BiocompatibleTransducers" filed Feb. 14, 1997, ("the 02323 PCT Patent Application")discloses improved implantable microactuators that are useful in thefully implantable hearing aid system disclosed in the 15087 PCT PatentApplication. The fully implantable hearing aid system disclosed in the15087 and 02323 PCT Patent Applications can operate for a period of fiveyears on a set of batteries, and produce sound levels of 110 dB. Thefully implantable hearing aid system described in these PCT PatentApplications is extremely compact, sturdy, rugged, and providessignificant progress towards addressing problems with presentlyavailable hearing aids.

As described in the 15087 PCT Patent Application, the microactuator ispreferably implanted into a fenestration that pierces the promontory ofthe cochlea, and uses stress-biased lead lanthanum zirconia titanate("PLZT") transducer material. Stress-biased PLZT materials exhibit veryhigh deflections and generate very high forces in comparison with otherexisting piezoelectric materials and/or structures. Such materialsprovide in a monolithic structure both a layer of conventional PLZT anda compositionally reduced layer from which the PLZT oxide has beenconverted to an electrically conductive Crete material. During operationof the transducer, the PLZT layer expands and contracts laterally uponapplication of an alternating current ("AC") voltage across the disk.Expansion and contraction of the PLZT layer flexes the diskback-and-forth due to differential expansion between the PLZT layer andthe unexpanding Crete layer. However, the Crete layer in that transducermaterial includes a metallic form of lead ("Pb") as one of itsconstituent elements.

Microactuators disclosed in the 15087 PCT Patent Application include amembrane diaphragm that provides good biological isolation for thetransducer. Moreover, use of the membrane diaphragm fully preserves, ormay, through the use of hydraulic amplification, actually enhancetransducer performance by magnifying the transducer's deflection ordisplacement. Although the transducers disclosed in the 15087 PCT PatentApplication usually attach the Crete layer to the membrane diaphragm andfully enclose that layer, a possibility still exists that lead may leachfrom the transducer.

SUMMARY OF THE INVENTION

An object of the present invention is to provide a fully biocompatibletransducer.

Another object of the present invention is to provide a fullybiocompatible transducer that is suitable for use in a microactuatorincluded in an implantable hearing aid system.

Another object of the present invention is to provide an transducer fora microactuator that is fabricated from only biocompatible materials.

Yet another object of the present invention is to provide a transducerfor a microactuator that is non-toxic.

Briefly, the present invention is a biocompatible, implantablemicroactuator adapted for inclusion in a fully implantable hearing aidsystem. The microactuator includes a hollow body having an open firstend and, in a preferred embodiment, open first and second faces that areseparated from the first end. The microactuator includes a firstflexible diaphragm that seals the body's first end, and that is adaptedfor deflecting outward from and inward toward the body. A secondflexible diaphragm seals across the body's first face, and a thirdflexible diaphragm seals across the body's second face therebyhermetically sealing the body. An incompressible liquid fills thehermetically sealed body.

In a preferred embodiment first and second biocompatible, unimorphs aremechanically coupled to and an integral part of respectively the secondand third flexible diaphragms. The first and second unimorphs areadapted for receiving an electrical driving signal upon the applicationof which the first and second unimorphs directly deflect respectivelythe second and third flexible diaphragms. Deflections of the second andthird flexible diaphragms are coupled by the liquid within the body tothe first flexible diaphragm.

Both the first and second unimorph include a plate of biocompatiblepiezoelectric material, preferably a lead zirconia titanate ("PZT") orPLZT material, to which is bonded a layer of biocompatible metal,preferably titanium, nickel, platinum, rhodium, palladium, gold, or ashape memory alloy such as nickel-titanium Naval Ordinance Laboratory("Nitinol"). The layer of biocompatible metal may be processed to applya stress-bias to the plate of biocompatible piezoelectric material. Thefirst and second unimorphs also include a thin, biocompatible electrodeapplied to each of the plates of biocompatible piezoelectric materialopposite the respective layers of biocompatible metal. Application ofthe electrical driving signal across the layer of biocompatible metaland the biocompatible electrode causes both of the unimorphs to deflect.To effect a stress-bias of the unimorphs, the unimorphs are annealed ata high temperature such as 500° C. to relieve all uncontrolled stress,and, in the instance of the shape memory alloy Nitinol, to establish ahard austenitic phase of the material. Cooling the unimorphs from theelevated temperature applies the requisite stress-bias to the unimorph.

These and other features, objects and advantages will be understood orapparent to those of ordinary skill in the art from the followingdetailed description of the preferred embodiment as illustrated in thevarious drawing figures.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a cross-sectional view illustrating a stress-biased PLZTtransducer in accordance with the 15087 PCT Patent Application;

FIG. 2 is a cross-sectional view illustrating a stress-biased unimorphtransducer in accordance with the present invention that has improvedbiocompatibility;

FIG. 3 is a phase diagram for a shape memory Nitinol alloy;

FIG. 4 is a cross-sectional view illustrating a stress-biased unimorphtransducer in accordance with the present invention that has improvedbiocompatibility;

FIG. 5 is a graph depicting deflection sensitivity of a stress-biasedunimorph transducer, such as those depicted in FIGS. 2 and 4, forvarious thicknesses of a metal layer that is bonded to a piezoelectricplate;

FIG. 6A is a partially sectioned elevational view of a microactuator fora fully implantable hearing aid system; and

FIG. 6B is a cross-sectional elevational view of the microactuator takenalong the line 6B--6B in FIG. 6A.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

FIG. 1 is a cross-sectional elevational view depicting a transducer 22as disclosed in the 15087 PCT Patent Application. The 15087 PCT PatentApplication discloses that the transducer 22 is preferably fabricatedfrom a thin circular disk of stress-biased PLZT material. This materialmay be one of the materials manufactured by Aura Ceramics and sold underthe "Rainbow" product designation. These PLZT unimorphs provide amonolithic structure one side of which is a layer 22a of conventionalPLZT material. The other side of the PLZT unimorph is a compositionallyreduced layer formed by chemically reducing the oxides in the nativePLZT material to produce a conductive Crete layer 22b. The conductivecermet layer 22b typically comprises about 30% of the total diskthickness, and typically may have a composition that contains up to 60%lead. Removing of the oxide from one side of the unimorph shrinks theconductive Crete layer 22b which bends the whole disk and puts the PLZTlayer 22a under compression. The PLZT layer 22a is therefore convexwhile the conductive Crete layer 22b is concave.

For use in the fully implantable hearing aid system, as depicted in FIG.1 the PLZT layer 22a and the conductive Crete layer 22b are respectivelyovercoated with thin metal layers to provide a PLZT electrode 24 and aCrete electrode 26. The electrodes 24 and 26 may be applied to thetransducer 22 in various different ways such as plating, evaporation,metal spraying etc. Application of a voltage across the electrodes 24and 26 causes the disk to become either more or less bowed, dependingupon the polarity of the applied potential.

Although in the fully implantable hearing aid system's microactuatordisclosed in the 15087 PCT Patent Applications the transducer 22 doesnot contact a subject's body, and may be hermetically sealed within themicroactuator, some of the properties of the metallic form of lead inthe conductive Crete layer 22b may be undesirable.

FIG. 2 illustrates an alternative structure for the transducer firstproposed in the 15087 PCT Patent Application that eliminates themetallic form of lead containing Crete conductive Crete layer 22b. InFIG. 2, a metal laminated unimorph 32 consists of a plate 34 ofbiocompatible piezoelectric material, such as a biocompatible PLZT orbiocompatible PZT, onto which is deposited a conductive metallic layer36. For biocompatible PLZT and PZT materials, lead occurs in an oxideform which appears to be non-toxic.

The 15087 PCT Patent Application discloses that in fabricating thelaminated unimorph 32 for the fully implantable hearing aid system thepiezoelectric plate 34 is lapped down to a thickness from 1.0 to 6.0mils, and then coated with a thin chromium layer 38 onto which is plateda thin nickel layer 36. The thin nickel layer 36 stresses thepiezoelectric plate 34 thereby mimicking the stress-bias of theconductive Crete layer 22b in the PLZT unimorph transducer 22. A thin,biocompatible metal electrode 44 is applied to the plate 34 opposite thelayer 36. In practice it has been found that the stress applied with aplated nickel layer 36 is usually low, and that it is difficult tocontrol the stress-bias because of the changing parameters in theplating bath.

The 15087 PCT Patent Application also proposes that a metal laminatedunimorph 32 may be fabricated by applying a thin layer 36 of a shapememory alloy, such as 5 to 20 microns of Nitinol,nickel-titanium-copper, or copper-zinc-aluminum, to the piezoelectricplate 34. After a layer 36 of such material has been applied to thepiezoelectric plate 34, heating or cooling the shape memory alloyestablishes a phase in which the alloy layer 36 applies compressive ortensile stress to the plate 34. As is apparent to those skilled in theart of shape memory alloys, hysteresis in a phase transition of thealloy maintains that stress upon removal of the heating or cooling.However, it has been determined experimentally that applying a shapememory alloy layer 36 to the plate 34 as described without a carefulthermal procedure does not necessarily yield a satisfactory structure.

The stress-bias required for the fully implantable hearing aid system'stransducer can be simply obtained through differential thermal expansionof materials deliberately created during the fabrication of theunimorph. Since in general for the same temperature change metals expandmore than ceramic dielectric materials, a ceramic layer may be easilysubjected to compressive stresses similar to those in Aura Ceramics'Rainbow products. However, the bond between the metal and the ceramicdielectric should be as thin and strong as possible.

The transducers required for the fully implantable hearing aid systemare fairly thin. The desired overall transducer thickness ranges from1.0 to 6.0 mils, providing a compromise between good deflection (forthin transducers) and ease of fabrication and ruggedness with thicktransducers. Consequently, it is possible to produce suitablestress-biased structures by vacuum deposition such as sputtering,evaporating or spraying metal onto a hot piezoelectric substrate,thereby producing a suitable metal phase. Upon cooling, greatercontraction of the metal layer in comparison with the piezoelectriclayer creates the required stress-bias. In principal, such a processyields a type of stress-bias similar to that produced by reduction andfabrication of Aura Ceramics' Rainbow products.

It is believed that metals with physical properties similar to those ofthe Rainbow material's conductive Crete layer 22b will yield nearoptimum performance for the fully implantable hearing aid system'stransducer. The thermal expansion of most piezoelectric materials isabout 2-4 ppm/° C., while the Crete expansion coefficient is about 10ppm/° C. Hence it is desirable to find biocompatible metals which havean expansion coefficient similar to that of the conductive Crete layer22b, and that have suitable mechanical strength. Materials having suchcharacteristics include Nitinol, titanium, platinum, rhodium, palladium,gold, and nickel, if suitably prepared. The first two materials, i.e.Nitinol and titanium, have excellent biocompatible properties. Both thethermal expansion coefficient (11 ppm/° C.) and the elastic modulus(8×10¹¹ Pa) of the austenitic phase of Nitinol match closely thecorresponding properties of the conductive Crete layer 22b of AuraCeramics' Rainbow products (respectively 10 ppm/° C. and 7×10¹¹ Pa).Therefore, for the same thickness of the conductive layers 22b and 36,and for the dielectric layers 22a and 34, it is reasonable to expectthat the same stress-bias will develop upon cooling since the differencebetween the transducers' processing temperature and operatingtemperature are approximately the same.

FIG. 3 depicts a thermal expansion diagram for Nitinol. Upon coolingfrom high temperature along a line 52 in FIG. 3 to a temperature T1,Nitinol exhibits a contraction coefficient of 11 ppm/° C. Along the line52 of the phase diagram, Nitinol exists as a harder austenitic phase.Upon reaching the temperature T1 (around 80° C.), a phase changecommences that converts Nitinol's hard austenitic phase material to asofter martensitic phase. Conversion to the martensitic phase duringcooling of Nitinol from the temperature T1 to a temperature T2 along aline 54 causes Nitinol to expand significantly, i.e. up to 0.5% of itslength. Linear expansion during Nitinol's phase transformation from theaustenitic phase to the martensitic form may be as much as 2000 ppm,equivalent to expansion of the material over almost 200° C. if the phasechange did not occur. At T2 (around 75° C.), Nitinol completes itsconversion to the martensite phase. Further cooling of Nitinol below T2along a line 56 in FIG. 3, Nitinol exhibits the martensitic contraction(expansion) coefficient of 6 ppm/° C. Nitinol's martensitic phase has anelastic modulus which is 2 to 3 times lower than the austenitic elasticmodulus, and the martensitic phase's yield strength is from 3 to 5 timesless than the austenitic phase.

Heating Nitinol along the line 56 from below the martensitic transitiontemperature T2 produces an expansion of the material up to a temperatureT3 (about 90° C.). Further heating of the material above the temperatureT3 along a line 58 commences conversion of Nitinol's softer martensiticphase material to the harder austenitic phase. Conversion to theaustenitic phase during heating of Nitinol from the temperature T3 to atemperature T4 along a line 58 causes, Nitinol to contractsignificantly, i.e. up to 0.5% of its length. At T4 (around 95° C.),Nitinol completes its conversion to the austenitic phase. Thetemperature hysteresis in Nitinol's phase change cycle is approximatelyof 10-20° C., but can be larger.

Because Nitinol's martensitic phase is much softer than the material'saustenitic phase, use of the material in the martensitic phase to applya pre-established stress-bias to the plate 34 depicted in FIG. 2requires a substantially thicker layer 36 of Nitinol than if thematerial is in the austenitic phase. Use of a thick Nitinol layer 36 ismore expensive, and is undesirable for various other reasons. Also, ifstress-bias is applied to the plate 34 by thermal expansion, asignificant portion of the desired stresses will be released if Nitinolundergoes the phase transition to the martensitic form.

Consequently, using a shape memory allow which has a phase changetemperature well below the operating temperature of the transducer, i.e.36° C., or, in general, room temperature, is preferable for a transducerto be included in the fully implantable hearing aid system. In the bulkmaterial, lowering the martensitic phase transition temperature T1 ofNitinol may be accomplished by adding small percentages of copper, iron,cobalt or chromium to the nickel-titanium alloy. However, for the fullyimplantable hearing aid system's transducer, the layer 36 is oftenapplied from the vapor phase, as it is difficult to find nickel-titaniumalloy foils having the required thickness. (However, it appears thatfoils having the requisite thickness may be obtained on special orderfrom Shape Memory Applications, Inc. of Santa Clara, Calif.)

The martensitic phase transition temperature T1 of sputterednickel-titanium films is typically approximately 60-80° C., which is toohigh for the fully implantable hearing aid system's transducer. It hasbeen observed that adding impurities inadvertently (as for examplecontaminants present on the surface of a poorly cleaned sample) maylower the martensitic phase transition temperature T1 significantly,without greatly affecting Nitinol's mechanical properties.Alternatively, it has been observed that co-sputtering nickel-titaniumtogether with metallic impurities, such as copper, iron, cobalt orchromium, typically with all constituent elements mixed in thesputtering target, may produce a martensitic phase transitiontemperature T1 well below room temperature. Furthermore, in manysputtering systems it has also been observed that depositing thenickel-titanium layer 36 with a large anode-cathode separation produce alower martensitic phase transition temperature T1.

In practice then, Nitinol material suitable for the transducer used inthe fully implantable hearing aid system may be prepared as follows.First, the nickel-titanium alloy including the copper, iron, cobalt orchromium impurity is sputtered from the target unto the plate 34, whichmay be at moderate temperature (e.g. 100-200° C.). Typically thesputtered Nitinol material will be amorphous. The laminated unimorph 32is then heated in an inert atmosphere or vacuum to 500° C. forapproximately one-half (0.5) hour to form the desired high strengthaustenitic phase. Heating the laminated unimorph 32 to 500° C. alsoremoves all stresses, and sets the temperature range over which thematerial will be stressed by differential thermal contraction.Alternatively, the nickel-titanium alloy including the copper, iron,cobalt or chromium impurity may be deposited directly on the plate 34 athigh temperature. A typical sputtering rate may be 6 minutes per micronof the Nitinol layer 36, with 2 kW of power, at a pressure of 2×10⁻³micron of argon. The austenitic phase is preserved during cooling toroom temperature. For Nitinol in the austenitic phase, the optimumthickness for the layer 36 is about 35 microns for a 75 microns thickplate 34, i.e. about 0.5. For a 50 micron thick PLZT plate 34, theoptimum thickness for the Nitinol layer 36 is about 23 microns.

As an alternative to Nitinol, sputtered films of either pure nickel orpure titanium that are annealed at 500° C. and cooled can also be usedvery advantageously for the fully implantable hearing aid system'stransducer, although nickel is not quite as biologically inert astitanium. Nickel's large coefficient of expansion (13 ppm/° C.) permitsa thinner layer 36 to produce the desired compression of the ceramiclaminated unimorph 32. Titanium's coefficient of expansion (9 ppm/° C.)closely matches that of the conductive Crete layer 22b. The coefficientsof expansion of other metals such as platinum, rhodium, palladium, orgold are also similar to that of the conductive Crete layer 22b.

For nickel, the optimum thickness of the layer 36 is approximately 0.26of the thickness of the plate 34, and the bowing of the laminatedunimorph 32 is about 20% larger at the same voltage than with Nitinol.For titanium, the optimum thickness of the layer 36 is about 0.40 of thethickness of the plate 34, and the bowing of the laminated unimorph 32is about 10% greater at the same voltage than for Nitinol. As comparedto the PLZT transducer 22, the optimum nickel structure provides up to a30% improvement for the same thickness of the ceramic plate 34 and thesame applied voltage.

The temperature difference, from the annealing temperature to roomtemperature, over which the differential contraction occurs togetherwith the difference in the coefficients of thermal expansion between theplate 34 and the layer 36 applies a stress-bias to the laminatedunimorph 32. Moreover, because the layer 36 may be appliedincrementally, it is therefore possible to establish a particularstress-bias for the laminated unimorph 32 using the method describedabove, and to then subsequently increase the thickness of the layer 36at a lower temperature, without significantly changing the existingstress-bias. The ability to establish a particular stress-bias in thepiezoelectric plate 34 independently of the total thickness of the layer36 permits establishing a thickness for the layer 36 which producesoptimum deflection of the laminated unimorph 32 in response to a voltageapplied across the layer 36 and the electrode 44. Such tailoring of theconductive Crete layer 22b is impossible for the Aura Ceramics' Rainbowproducts.

FIG. 4 depicts a stress-biased laminated unimorph 62 in accordance withthe present invention. The laminated unimorph 62 typically includes a 75micron (3 mils) thick piezoelectric plate 64 preferably of abiocompatible ceramic PZT or PLZT material. Using titanium for thestress-biasing material, the optimum thickness for the metal layer 66-68is 30 microns, although a 20 micron thickness produces 95% thedeflection of a 30 micron thick layer 66-68. Thus, for example, a 10micron thick metallic layer 66 of titanium (or any suitable thickness toprovide the desired stress-bias) is first sputtered onto the laminatedunimorph 62, annealed at 500° C. as described above, and then cooled.Subsequently, another metallic layer 68, e.g. 10 microns or any otherappropriate thickness of titanium or any other suitable metal, can thenbe sputtered or bonded onto the layer 66 without annealing to increasethe total thickness of the metal layers 66 and 68 to that which providesoptimum deformation characteristics for the laminated unimorph 62. Thedeposition or bonding of the layer 68 does not significantly change thestress-bias applied by the layer 66, although insignificant stresses maysometimes result from the deposition of the layer 68. A thin,biocompatible metal electrode 72 is applied to the plate 64 opposite thelayers 66 and 68.

The layers 66 and 68 of FIG. 4 may be of different materials, if sodesired. For example the first layer 66 may be titanium, providingexcellent adhesion to the plate 64, while the second layer 68 may benickel which sticks well to titanium. Vice versa, the first layer 66 maybe nickel to create a particular stress-bias with a relatively thinlayer of metal. The nickel layer 66 may then be subsequently coated withtitanium to improve the biocompatibility of the laminated unimorph 62.Typically, the laminated unimorph 62 has a total thickness from 1.0 to6.0 mils, with the combined thickness of the layers 66 and 68 equalingfrom 0.5 to 0.15 the thickness of the plate 64. Since it is difficult tocut a bowed material, often it is desirable to shape the plate 64 to itsdesired size as a flat sheet, and then deposit the metal layers 66 and68. Note that nickel can be readily plated in various forms (e.g.electrolytic or electroless), to the required thickness. Very thinlayers of chromium (a few hundred Angstroms) can be applied to the plate64 if necessary to improve adhesion before depositing the layer 66.

Other materials which may be used for the transducer of the fullyimplantable hearing aid consist of a PLZT or PZT plate that isphysically bonded to a metal sheet at high temperature. Here both aceramic plate and a metal sheet are heated to the same high temperature,with a solder or other suitable bonding material disposed between theceramic plate and the metal sheet. Upon cooling, the unimorph thuscreated exhibits the desired stress-bias. The biocompatible materialused for the metal sheet include titanium, nickel, platinum, rhodium,palladium, gold, or Nitinol foils, about 0.5 to 3.5 mils thick, in thesame metal/ceramic thickness ratio as described above. To avoid possibleleaching of the metallic form of lead, suitable bonding materialsinclude various types of lead free solder, for example indium alloysparticularly in paste form. Such solder pastes may be screen printedunto the metal sheet or the piezoelectric ceramic plate in very thinlayers e.g. 5 microns thick. The metal sheet thickness is about 0.5 to0.15 of the thickness of the piezoelectric ceramic plate. Suitablebonding temperatures are 150-400° C. Typically a weight is put on thestacked ceramic plate and metal sheet to press them together duringbonding. The weight may be removed during cooling. Usually thepiezoelectric ceramic plate will first be coated with a very thin layerof silver paste that is then fired onto the piezoelectric ceramic athigh temperature. An electrically conductive metal electrode on theceramic side, opposite the bonded metal sheet, should be as thin aspossible, to avoid creating stresses which oppose the stresses generatedby the bonded metal sheet. These stress-biased structures, 1.25 inch indiameter, 15 mils thick produce axial displacements of 100 micron inresponse to an applied signal of 150 volts. Hence for a 4 mil thickstructure, 2.5 mm in diameter, with 10 V drive, the expected axialtransducer displacement is about 0.6 micron, more than adequate for thetransducer included in the microactuator of the fully implantablehearing aid system.

FIG. 5 depicts generally deflection sensitivity of a laminated unimorph,such as the unimorphs 32 and 62, as a function of the thickness of themetal layer 36 or 66-68. A graph 76, indicated by "+" symbols, in FIG. 5depicts deflection sensitivity versus thickness of the metal layer 36 or66-68 for a 2.0 mm square or disk-shaped plate 34 or 64 of 3203 PZTceramic piezoelectric material that is 75.0 microns thick computed usinga formula by Timoshenko for bimetallic springs, Journal Optical Societyof America, vol. 11, no. 233, 1925. A line graph 78 in FIG. 5 depictsdeflection sensitivity versus thickness of the metal layer 36 or 66-68for the same piezoelectric material computed using a formula by Chu etal., J. Micromech. Microeng. 3 (1993), also for bimetallic springs. Asdepicted by the graphs 76 and 78, increasing the thickness of the metallayer 36 or 66-68 initially increases the deflection sensitivity of theunimorph 32 or 62. Moreover, the deflection sensitivity of the unimorph32 or 62 continues to increase with increasing thickness of the metallayer 36 or 66-68 until reaching a maximum value that extends across abroad range of thicknesses.

Heating the laminated unimorph 32 or 62 to 500° C. exceeds the Curiepoint of the piezoelectric ceramic material. Therefore the piezoelectricceramic material needs to be repoled preferably as the laminatedunimorph 32 or 62 cools. Alternatively the laminated unimorph 32 or 62can be repoled after cooling. PZT materials suitable for the laminatedunimorph 32 or 62 are identified by various commercial names such asPZT-4, PZT-5A, PZT-5H, PZT-8, C3100, and C3200, and in particular 3203,3199 or 3211, the latter materials being manufactured by Motorola, Inc.Desirable properties for the laminated unimorph 32 or 62 include a veryhigh value for the piezoelectric constant d31, which determinestransverse contraction, and for good mechanical machinability. The 3203material appears best in these respects, and is the presently preferredmaterial.

FIGS. 6A and 6B depict a microactuator 82 adapted for inclusion in afully implantable hearing aid system. The microactuator 82 includes ahollow body 84, depicted only in FIG. 6B, from one end of which projectsan L-shaped, flanged nozzle 86. The flanged nozzle 86 has an open firstend 88 that is sealed by a flexible diaphragm 92 that may be deflectedoutward from and inward toward the flanged nozzle 86. As described ingreater detail in the 02323 PCT Patent Application, the first end 88 andthe diaphragm 92 are adapted for implantation into a fenestration formedthrough a promontory that is located between a subject's middle andinner ear. The body 84 has two open faces 94a and 94b that are separatedfrom the first end 88. Each of the faces 94a and 94b are respectivelysealed by flexible diaphragms 96a and 96b which, in combination with thediaphragm 92, hermetically seal the body 84. As depicted in FIG. 6B, thediaphragms 96a and 96b respectively have cross-sectional areas that arelarger than a cross-sectional area of the diaphragm 92. While thepreceding description of the body 84 identifies various individual partsthereof, the body 84 may, in fact, be provided by a one-piece can formedfrom a material suitable for the diaphragms 96a and 96b.

The hermetically sealed hollow body 84 is filled with the incompressibleliquid 98. Respectively secured to each of the diaphragms 96a and 96bare transducers 102 which face each other. In accordance with thepresent invention, the transducers 102 of the microactuator 82 areprovided by the laminated unimorphs 32 or 62. Each of the transducers102 are electrically connected to a miniature cable 104 to expand orcontract in opposite direction toward or away from each other inresponse to the same voltage applied across each of the transducers 102.This driving motion of the transducers 102 applied to the diaphragms 96aand 96b forces the liquid 98 toward or away from the diaphragm 92 thatis located in a subject's inner ear thereby deflecting the diaphragm 92.While the microactuator 82 preferably employs a pair of transducers 102,a microactuator 82 in accordance with the present invention may haveonly a single transducer 102, or each transducer 102 of the pair mayhave a different shape and/or size.

As described in greater detail above, the transducer 102 exhibitsmaximum deflection sensitivity when the combined thicknesses of themetal layer 36 or 66-68 together with that of the diaphragm 96a or 96bis within the broad range of optimum thickness computed using thetheories of Timoshenko and/or Chu. Thus, the laminated unimorph 62 maybe built by first depositing the layer 66, and by then bonding the layer66 to the diaphragm 96a or 96b which therefore provides the second layer68. One method for assembling the laminated unimorph 62 is by firstbonding the layer 66 to the diaphragm 96a or 96b, which is then securedto the remainder of the body 84 by electron-beam or laser welding.Alternatively, the diaphragm 96a or 96b may first be welded onto theremainder of the body 84, after which the layer 66 is bonded to thediaphragm 96a or 96b.

An alternative stress-biased unimorph which may be used for thetransducer 102 is a NASA Langley Research Center's THin-layer compositeUNimorph ferroelectric DrivER and sensor ("THUNDER") high-displacementactuator. As described in NASA publications, the THUNDER transducerconsists of a piezoelectric plate intimately attached to a pre-stressinglayer with LARC-SI, a NASA Langley Research Center developed polyimideadhesive. Reportedly, the pre-stressing layer, provided by adhesivebacked foils, is thermally bonded together and to the piezoelectricplate in vacuum. This initial vacuum and thermal processing yields apartially pre-stressed THUNDER unimorph. Fabrication of the THUNDERunimorph is then completed by pressing the partially pre-stressedunimorph onto a curved fixture while heating it to increase theunimorph's pre-stress.

To isolate the transducers 102 from a subject's body, the body 84 andthe transducers 102 of the microactuator 82 are preferably enclosedwithin a hermetically sealed titanium housing 112. As illustrated inFIG. 6B the housing 112 is joined to the flanged nozzle 86 around aperimeter of a flange 114 such as by electronbeam or laser welding.Alternatively, the body 84 and the transducers 102 may be enclosed witha with parylene coating thereby isolating them from a subject's body.

Anatomical considerations permit the transducers 102 to extend aconsiderable distance into a subject's middle ear cavity, and alsopermit shapes for the body 84 and the transducers 102 that differ fromthose depicted in FIGS. 6A and 6B. The base of the body 84 adjacent tothe flanged nozzle 86 can be very narrow and the length of the body 84and transducers 102 extending outward from the flanged nozzle 86enlarged so that the volume of the liquid 98 displaced by thetransducers 102 becomes quite large. In this way, the faces 94a and 94band the transducers 102 can be shaped, twisted and tilted to fit asubject's middle ear cavity, and are not restricted to the spaceavailable locally at the implantation site.

While the illustrations of FIGS. 6A and 6B depict the diaphragms 96a and96b as being oriented parallel to the diaphragm 92 with the diaphragms96a and 96b parallel to each other, other orientations of the diaphragms96a and 96b with the respect to the diaphragm 92 are within the scope ofthe present invention. Accordingly, the diaphragms 96a and 96b can beoriented at a skewed angle with respect to the flanged nozzle 86 anddiaphragm 92 to prevent the transducers 102 from interfering with anossicular chain 21 or other structures located within a subject's middleear. The flanged nozzle 86 provides good anchoring to the promontorywithout requiring extra room which would otherwise reduce spaceavailable for the transducers 102.

Although the present invention has been described in terms of thepresently preferred embodiment, it is to be understood that suchdisclosure is purely illustrative and is not to be interpreted aslimiting. While it appears that stress-biased transducers 102 offersuperior performance for the microactuator 82, it also appears thattransducers 102 that are not stress-biased offer performance adequatefor the fully implantable hearing aid system. Consequently, withoutdeparting from the spirit and scope of the invention, variousalterations, modifications, and/or alternative applications of theinvention will, no doubt, be suggested to those skilled in the art afterhaving read the preceding disclosure. Accordingly, it is intended thatthe following claims be interpreted as encompassing all alterations,modifications, or alternative applications as fall within the truespirit and scope of the invention.

What is claimed is:
 1. A biocompatible unimorph adapted for use as atransducer of a microactuator included in a fully implantable hearingaid system, the unimorph comprising:a plate of biocompatiblepiezoelectric material; a layer of biocompatible metal deposited ontosaid plate of biocompatible piezoelectric material; and a thin,biocompatible electrode applied to said plate of biocompatiblepiezoelectric material opposite said layer of biocompatible metal,whereby application of an electric potential across said layer ofbiocompatible metal and said biocompatible electrode causes the unimorphto deflect.
 2. The biocompatible unimorph of claim 1 wherein said plateof biocompatible piezoelectric material is formed from lead zirconiatitanate ("PZT") material.
 3. The biocompatible unimorph of claim 1wherein said plate of biocompatible piezoelectric material is formedfrom lead lanthanum zirconia titanate ("PLZT") material.
 4. Thebiocompatible unimorph of claim 1 wherein said layer of biocompatiblemetal is formed from a material chosen from a group consisting oftitanium, platinum, rhodium, palladium, gold, and nickel.
 5. Thebiocompatible unimorph of claim 1 herein said layer of biocompatiblemetal is formed from a shape memory material.
 6. The biocompatibleunimorph of claim 5 wherein the shape memory material forming said layerof biocompatible metal is a nickel-titanium Naval Ordinance Laboratory("Nitinol") alloy.
 7. The biocompatible unimorph of claim 1 wherein:saidlayer of biocompatible metal is deposited onto said plate ofbiocompatible piezoelectric material; andwherein, after depositing saidlayer of biocompatible metal onto said plate of biocompatiblepiezoelectric material, the unimorph is: heated to a temperature whichrelieves stress in the unimorph; and cooled whereby said layer ofbiocompatible metal appliesa stress-bias to said plate of biocompatiblepiezoelectric material.
 8. The biocompatible unimorph of claim 7 whereinthe material forming said layer of biocompatible metal is chosen from agroup consisting of titanium, nickel, platinum, rhodium, palladium,gold, and Nitinol; and wherein following depositing the unimorph isheated to approximately 500° C.
 9. The biocompatible unimorph of claim 1wherein:said layer of biocompatible metal is deposited onto said plateof biocompatible piezoelectric material that is heated to a temperaturewhich relieves stress in the unimorph, material that forms said layer ofbiocompatible metal; andwherein, after depositing said layer ofbiocompatible metal onto said heated plate of biocompatiblepiezoelectric material, the unimorph is: cooled whereby said layer ofbiocompatible metal applies a stress-bias to said plate of biocompatiblepiezoelectric material.
 10. The biocompatible unimorph of claim 9wherein during depositing said plate of biocompatible piezoelectricmaterial is heated to approximately 500° C., and wherein the materialforming said layer of biocompatible metal is chosen from a groupconsisting of titanium, nickel, platinum, rhodium, palladium, gold, andNitinol.
 11. The biocompatible unimorph of claim 1 wherein said plate ofbiocompatible piezoelectric material has a thickness, and said layer ofbiocompatible metal has a thickness that is between 0.15 and 0.5 timesthe thickness of said plate.
 12. The biocompatible unimorph of claim 1wherein:a fraction of said layer of biocompatible metal is firstdeposited onto said plate of biocompatible piezoelectric material; theunimorph is cooled whereby said layer of biocompatible metal thendeposited onto said plate of biocompatible piezoelectric materialapplies a stress-bias to said plate; and additional material for formingsaid layer of biocompatible metal is bonded to said existing layer ofbiocompatible metal to complete formation of said layer of biocompatiblemetal.
 13. The biocompatible unimorph of claim 12 wherein said layer ofbiocompatible metal first deposited onto said plate of biocompatiblepiezoelectric material together with the additional material for formingsaid layer of biocompatible metal bonded to said existing layer ofbiocompatible metal have a combined thickness which provides optimumdeflection sensitivity for the unimorph.